The background description provided herein is for the purpose of generally presenting the context of the disclosure. Work of the presently named inventors, to the extent the work is described in this background section, as well as aspects of the description that may not otherwise qualify as prior art at the time of filing, are neither expressly nor impliedly admitted as prior art against the present disclosure.
In positron emission tomography (PET) imaging, a tracer agent is introduced into the patient, and the physical and bio-molecular properties of the agent cause it to concentrate at specific locations in the patient's body. The tracer emits positrons, resulting in an annihilation event occurs when the positron collides with an electron that produces two gamma rays (at 511 keV) traveling at substantially 180 degrees apart.
PET imaging systems use detectors positioned around the patient to detect coincidence pairs of gamma rays. A ring of detectors can be used in order to detect gamma rays coming from each angle. Thus, a PET scanner can be substantially cylindrical to be maximize the capture of the isotropic radiation. A PET scanner can be composed of several thousand individual crystals (e.g., Lutetium Orthosilicate (LYSO) or other scintillating crystal) which are arranged in two-dimensional scintillator arrays that are packaged in modules with photodetectors to measure the light pulses from respective scintillation events. For example, the light from respective elements of a scintillator crystal array can be shared among multiple photomultiplier tubes (PMTs) or can be detected by silicon photomultipliers (SiPMs) having a one-to-one correspondence with the elements of a scintillator crystal array.
When PMTs are used as the photodetectors, Anger logic can be used based on the relative geometry between the scintillating crystal elements and the respective PMTs, which determines the relative pulse energy measured by the photodetectors. Using Anger logic/arithmetic and a floodmap calibrated lookup table, the relative pulse energies of the PMTS are compared to determine at which position within the crystal array (i.e., which crystal element) the scintillation event occurred.
To reconstruct the spatio-temporal distribution of the tracer via tomographic reconstruction principles, each detected event is characterized for its energy (i.e., amount of light generated), its location, and its timing. By detecting the two gamma rays, and drawing a line between their locations, i.e., the line-of-response (LOR), one can determine the likely location of the original disintegration. The timing information can also be used to determine a statistical distribution along the LOR for the annihilation even based on a time-of-flight (TOF) between the two gamma rays. While this process will only identify a line of possible interaction, by accumulating a large number of those lines, and through a tomographic reconstruction process, the original distribution can be estimated.
Single-photon emission computed tomography (SPECT) is similar to PET except a collimator is used to restrict the solid angle of gamma rays incident on the respective detector elements (e.g., the respective elements in the scintillator crystal array), making reconstruction possible using single gamma ray detection events as opposed to requiring coincidences to determine a LOR.
Both PET and SPECT imaging depend on the ability to determine the position at which a gamma ray is detected. However, scatter events can result in a part of the gamma ray energy being deposited in the original detection crystal element with the scattered gamma ray depositing the remaining energy in one or more other crystal elements, generating ambiguity regarding which crystal element was the original detection element. Conventional methods of position correction to resolve suffer from at least two shortcomings. First, conventional position-correction methods are limited to single scatter events (i.e., the gamma ray energy scatters just once, resulting in the energy being deposited in only two crystal elements). Second, conventional position-correction methods fail when the two crystal elements have approximately equal energy from the gamma ray. Accordingly, more robust and scalable methods are desired for position correction/determination when the gamma ray is scattered/shared among multiple scintillator crystal elements in a gamma detector/camera.